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1. Field of the Invention
This invention relates to detector device that may be used to detect radiation emitted by positron annihilations.
2. Description of the Related Art
Positron emission tomography (PET) is a powerful imaging technique providing capabilities which no other technology is likely to match. Its most unique feature is its ability to image tissue differentiated by function. In particular, the field is currently dominated by use of the imaging agent 18-FDG, which is a glucose analog. This agent is capable of labeling tissue in proportion to the tissue""s ability to metabolize glucose. Such imaging shows great promise in the field of oncology since many tumors have altered metabolic processes relative to surrounding normal tissue. As a result, a large variety of primary and metastatic tumors are well visualized with the technique. No other technique is capable of distinguishing tissue according to its metabolic status.
Furthermore, PET produces three-dimensional imaging with unparalleled quantitative content. The basic high quality imaging characteristics of PET, including high sensitivity, attenuation correction, high contrast, and fully quantitative 3-dimensional character, enhance the potential of this technology even further. Finally, the chemistry of PET radiopharmaceuticals is currently in its infancy. It is very likely that, with wide availability of the imaging devices, even more powerful radiopharmaceuticals will follow.
Some of the promising applications of the 18-FDG radiotracer (half life=110 minutes) include imaging of brain tumors (DiChiro G. Positron emission tomography using [18-F]fluorodeoxyglucose in brain tumors: a powerful diagnostic and prognostic tool. Invest Radiol 1986; 22: 360-371; Coleman R E, Hoffinan H M, Hanson M W, Sostman H D, Schold S C. Clinical application of PET for the evaluation of brain tumors. J. Nucl. Med. 1991; 32:616-622), lung cancer (Patz E F, Lowe V, Hoffinan J M, et al. Evaluation of focal pulmonary abnormalities with [18-F]-fluoro-1-deoxyglucose and positron emission tomography. Radiology 1993; 188: 487-490.; Gambhir S S, Hoh C K, Phelps M E, Madar I, Maddahi J. Decision tree sensitivity analysis for cost-effectiveness of FDG-PET in the staging and management of non-small-cell lung carcinoma. J Nucl Med 1996; 37: 1428-1436; Patz E F, Lowe V J, Hoffiman J M, Paine S S, Harris L K, Goodman P C. Persistent or recurrent brochogenic carcinoma: detection with PET and 2-[18-F]-2-deoxy-D, glucosee. Radiology 1994; 191: 379-382), melanoma, lymphoma, colorectal carcinoma, breast cancer, head and neck cancer, gynecologic malignancy and bone and soft tissue malignancy (Conti P S, Lilien D L, Hawley K, Kepper J, Grafton S T, Bading J R. PET and [18-F]-FDG in oncology: a clinical update. Nucl Med Biol 1996;23: 717-735). Regional distribution centers are actively being established in preparation for broad usage of the 18-FDG radiopharmaceutical.
Unfortunately, PET camera technology is poorly suited for many important oncology applications, in which large regions of the body frequently need to be imaged. The typical ring style PET cameras offer a limited 12-16 cm axial field of view. Therefore, for imaging of large longitudinal sections of the body, acquisition of many individual slice images with sequential translation of the patient along the axis of the camera is required. These procedures can and frequently do require several hours of imaging, when large areas of the body are scanned for metastatic disease, for example. Clearly, such multiple view imaging also reduces sensitivity within any given imaging slice due to the limited time available for each region. The limited field of view of current PET detectors is dictated primarily by the high cost of the large array of crystals and photomultiplier tubes (PMT). In PET imaging, positron annihilation photons at 511 keV are imaged in coincidence, classically with large arrays of crystal detectors. These detectors must encompass a large portion of the 360 degree geometry around the patient. They must also be capable of stopping the 511 keV gammas with high efficiency. Furthermore, excellent time resolution is required because of the high singles rate and the large number of crystals needed. Because of this combination of requirements, classical PET cameras utilize crystals of BGO (Bi4Ge3O12) (Moses W W, Derenzo S E, Budinger T F. PET detector modules based on novel detector technologies. Nucl Instr Meth 1994; A 353: 189-194). The high Z (mass number) of the material (principally bismuth) is advantageous for scatter-free detection of 511 keV by the photoelectric process. In NaI, over 75% of the interactions are compton scatter, which frequently results in multiple interaction points in the crystal array. If such scattered events are counted, many erroneous events result, and scatter is included. If they are not counted, than the upper limit on efficiency is less than 25%. By contrast, BGO provides a 43% photoelectron interaction fraction, yielding much cleaner and more efficient imaging. Crystal detectors of BGO are expensive and produce limited light output, driving up the cost of photomultiplier tubes. The base materials cost of assembled crystals and photomultipliers runs between $350,000 and $700,000 in today""s cameras, and as a result the largest longitudinal field of view available commercially is limited to 24 cm (Moses W W, Derenzo S E, Budinger T F. PET detector modules based on novel detector technologies. Nucl Instr Meth 1994; A 353: 189-194).
Because of this high fixed materials cost, totally assembled and supported cameras cost from $1.5 million to over $3 million. Largely because of these exceedingly high costs, only about 60 dedicated PET imaging centers are in operation in the U.S. The small number of facilities combined with the limitation of large field of view imaging means that a very small fraction of the diagnostic imaging need for oncology patients is being met with today""s technology.
In the hope of reducing high instrument cost and in order to increase the field of view, alternative systems are being actively pursued in the industry. Moses et al. have recently summarized the technical characteristics desired for high quality PET detectors (Moses W W, Derenzo S E, Budinger T F. PET detector modules based on novel detector technologies. Nucl Instr Meth 1994; A 353: 189-194). These requirements, listed in order of decreasing importance, are: high efficiency ( greater than 85%), high spatial resolution ( less than 5 mm FWHM), low cost ( less than $600/sq. inch), low dead time ( less than 4 xcexcsec/sq. inch), good timing resolution ( less than 5 nsec FWHM), and energy resolution ( less than 100 keV FWHM). The problem of identifying a lower cost PET detector has been exceptionally intractable. The largest effort has focused on application of large area single crystal NaI devices in paired geometry. Such dual imaging systems are widely available for emission tomography, and the hope in the nuclear medicine instrumentation industry is that such cameras could provide a lower cost alternative for at least limited work with 18-FDG. However, these systems have very severe drawbacks caused by the gender and geometry of their crystals. Although the 1 cm thick NaI crystal utilized by most of these systems is adequate for Tc-99m, it provides only a 5% photoelectron efficiency at 511 keV, driving the coincidence efficiency down to an extremely limiting 0.2%. This efficiency is more than two orders of magnitude less than that of the conventional BGO PET camera. As a result, such systems require extremely lengthy data acquisition times. Furthermore, because very high singles rates are present in the uncollimated geometry, these systems saturate at very low injected activity levels.
Another alternative that has been explored is the use of the multiwire proportional chamber (MWPC). It is widely recognized that multiwire systems employed together with 511 keV converter modules can provide PET images at substantially lower costs. Despite extensive development efforts, however, these systems have very limited detection efficiencies of 10-30%, and depending on the converter system utilized, have either poor coincidence time resolution of 88 nsec compared to 5 nsec for the BGO systems, or poor spatial resolution of 5-11 mm compared to less than 6 mm for the BGO systems (Wells K, Visvikis D, Ott R J. Performance of a BaF2-TMAE detector for use in PET. IEEE Trans Nucl Sci 1994; to be published; McKee B T A, Dickson A W, Howse D C. IEEE Trans Med Img. MI-13 (1994) 176).
A novel wire detector design is here proposed which in part is based on the modem straw tube technology being extensively employed in contemporary high energy physics. Such straw tube detectors can be very inexpensively produced to cover large volumes in close-packed arrays of 5 mm or smaller individual tube diameter. Because of the small size and use of modern, high-speed counting gases, these tubes can achieve very good time resolution of 10-25 nsec (FWHM). Given the very thin walls (0.001xe2x80x3-0.0015xe2x80x3) of such tubes, it is proposed to wrap them with thin lead converters or to incorporate such thin lead foil in the wall itself which can effectively convert the 511 keV radiation to energetic photoelectrons with sufficient energy to easily penetrate the tube wall. Such lead converter equipped tubes may be employed in an annular, close-packed array, with a sufficient number of tubes and therefore aggregate lead foil thickness to effectively stop 511 keV radiation. Readout of the converter wrapped tube struck by an event together with longitudinal position along the converter wrapped tube will allow position determination in three dimensional space of 5 mm FWHM or better. Such readout techniques have been well developed and proven in the high energy physics field. The estimated cost of such a detector array, based on documented costs of existing high energy physics systems scaled up to volume commercial production, is $25 per square inch of detector surface, or about 20 times less than the $600 per square inch for the BGO crystal array (Moses,1994). Through reduction of the lead foil thickness to a minimum providing efficient photoelectron escape into the thin-walled straw tube, high detection efficiency approaching that of BGO is feasible. Therefore, this design can provide a PET imaging system with excellent characteristics and with substantial reduction of cost. It would thereby be feasible to construct economical PET systems of far larger area, potentially encompassing the entire patient torso, which would be of great benefit for oncology applications utilizing 18-FDG.